JP6270074B2 - Processing conditions for producing bioabsorbable stents - Google Patents

Description

  The present invention relates to laser processing tubing to form a stent.

  The present invention relates to laser processing of devices such as stents. Laser machining refers to the removal of material that occurs through the interaction of the laser and the target material. Generally speaking, these processes include laser drilling, laser cutting, and laser grooving, marking, or scribing. The laser processing process transfers photon energy to the target material in the form of thermal energy or photochemical energy. Material is removed by melting and blowing or by direct vaporization / ablation.

  When the substrate is laser machined, energy is transferred to the substrate. As a result, the region beyond the cutting edge is altered by its energy, thereby affecting the properties within this region. In general, changes in properties are inconvenient for the proper functioning of the device being manufactured. Thus, it is generally desirable to reduce or eliminate energy transfer beyond the material being removed, thus reducing or eliminating the extent of modification and the size of the affected area.

  One of many medical applications for laser processing involves the creation of a radially expandable endoprosthesis that is configured to be implemented within a body lumen. An “internal prosthesis” corresponds to an artificial device placed inside the body. “Lumen” refers to a cavity of a tubular organ, such as a blood vessel.

  A stent is an example of such an endoprosthesis. Stents are generally cylindrical devices that function to keep open and possibly expand blood vessels or certain areas of other anatomical lumens such as the urinary tract and common bile duct. Stents are often used in the treatment of atherosclerotic stenosis in blood vessels. “Stenosis” refers to a narrowing of a body passage or opening, or a contraction of its diameter. In such treatment, the stent reinforces the body's blood vessels and prevents restenosis after angioplasty within the vasculature. “Restenosis” refers to the recurrence of stenosis in a blood vessel or heart valve after its apparently successful treatment (stenting or valvuloplasty according to balloon angioplasty).

  Treatment of affected sites or lesions with stents requires both delivery and deployment of the stent. “Delivery” refers to the introduction and transfer of a stent through a body lumen to an area within a blood vessel in need of treatment, such as a lesion. “Deployment” refers to expanding the stent within the lumen at the treatment area. Delivery and deployment of the stent involves positioning the stent around one end of the catheter, inserting that end of the catheter through the skin and into the body lumen and advancing the catheter to the desired treatment location within the body lumen. This is done by expanding the stent in position and removing the catheter from the lumen.

  In the case of a balloon expandable stent, the stent is mounted around a balloon placed on the catheter. Stent placement generally requires that the stent be pressed or crimped onto the balloon. The stent is then expanded by inflating the balloon. The balloon can then be deflated and the catheter can be withdrawn. In the case of a self-expanding stent, the stent may be secured to the catheter via a stretchable sheath or sock. When the stent is in the desired body position, the sheath can be withdrawn, thereby allowing the stent to self-expand.

  The stent must be able to meet several mechanical requirements. First, the stent must be able to withstand the structural load applied to the stent when supporting the vessel wall, i.e., radial compressive force. Therefore, the stent must have adequate radial strength and stiffness. Radial strength is the ability of a stent to resist radial compression forces. After expansion, the stent is properly sized and shaped over its useful life, despite the various forces that can cause it to compress, including periodic loads induced by the beating heart. Must be maintained. For example, a radial force can be prone to retract the stent inward. In general, it is desirable to minimize retraction.

  Furthermore, the stent must have sufficient flexibility to allow crimping, expansion, and cyclic loading. The stent should have sufficient resistance to failure so that stent performance is not adversely affected during crimping, expansion, and cyclic loading.

  Finally, the stent must be biocompatible so as not to induce any adverse vascular response.

  The structure of a stent is generally composed of scaffolding that includes a pattern or network of interconnecting structural elements often referred to in the art as struts or bar arms. Scaffolding can be formed from a wire, tube, or sheet of material rolled into a cylinder. The scaffolding is designed so that the stent can be radially compressed (so that it can be crimped) and radially expanded (so that it can be deployed).

  Stents are made of a number of materials such as metals and polymers, including biodegradable polymer materials. Biodegradable stents require that the stent be present in the body for a limited period until, for example, its intended function of achieving and maintaining vascular patency and / or drug delivery is achieved. This is desirable in a number of potential therapeutic applications.

  Stents can be made by forming a pattern on a tube or sheet using laser processing. The basic laser-material interaction is similar, but there are several aspects between the various materials (metal, plastic, glass, ceramics, etc.), i.e. various absorption characteristics. This is important in selecting the appropriate wavelength to produce the desired result. After the appropriate wavelength is selected, the combination of pulse energy and pulse duration defines the optimum process conditions for that type of material. The properties of biodegradable polymers such as PLLA and PLGA tend to be very sensitive to energy transfer, such as from laser processing. Laser parameters and laser-material interactions to help select laser systems and define processing parameters that allow faster laser processing of biodegradable stents and minimize adverse effects on their properties It takes a lot of effort to understand.

  Various embodiments of the present invention are methods for laser processing a substrate to form a stent, the method comprising providing a thin-walled polymer substrate and cutting through the substrate wall. Laser processing a thin-walled polymer substrate with a laser beam having a certain pulse width and wavelength to form a structural element having a processed edge surface, wherein the laser beam is adjacent to the processed edge surface Modifying the substrate within a surface area to a minor degree, the modification comprising voids, cracks, or combinations thereof that contribute to variations in the modulus of the polymer with distance from the edge surface, And selecting the pulse width and wavelength so that voids or cracks are present at a depth of 2 microns or less, or the elastic modulus converges at 4 microns or less.

  An additional embodiment of the present invention is a method of laser processing a substrate to form a stent, comprising providing a thin walled PLLA polymer substrate, and applying a laser beam to the thin walled PLLA polymer substrate. Laser machining and cutting through the wall to form a structural element having a machined edge surface, the pulse width and wavelength of the laser beam are in the green range and the pulse width is 1-10 ps Including methods.

  A further embodiment of the present invention provides a plurality of interconnected structural elements formed by laser machining a thin wall PLLA polymer substrate with a laser beam that cuts through the wall to form the structural elements. The structural element has a side wall corresponding to the processed edge surface, the surface area adjacent to the side wall has damage caused by the interaction of the laser beam with the substrate, A polymer stent body comprising voids or cracks dispersed within a surface region from the edge surface to a depth of 2 microns or less.

FIG. 1 is a diagram of a machine controlled system for laser machining a tube. FIG. FIG. 3 is an enlarged axial view of a region where a laser beam interacts with a tube. FIG. 2 is a diagram of a portion of a strut or structural element formed by laser processing a substrate. FIG. 6 is a cross-sectional view of a portion of a strut perpendicular to a processed edge surface. 3 is an SEM image of a surface region adjacent to a laser processed edge surface for a combination of pulse width and wavelength. 3 is an SEM image of a surface region adjacent to a laser processed edge surface for a combination of pulse width and wavelength. 3 is an SEM image of a surface region adjacent to a laser processed edge surface for a combination of pulse width and wavelength. 3 is an SEM image of a surface region adjacent to a laser processed edge surface for a combination of pulse width and wavelength. FIG. 6 shows the modulus of elasticity for the combination of pulse width and wavelength versus displacement into the surface of the laser processed edge. FIG. 6 shows the modulus of elasticity for the combination of pulse width and wavelength versus displacement into the surface of the laser processed edge. FIG. 6 shows the modulus of elasticity for the combination of pulse width and wavelength versus displacement into the surface of the laser processed edge. FIG. 6 shows the modulus of elasticity for the combination of pulse width and wavelength versus displacement into the surface of the laser processed edge. FIG. 6 shows the modulus of elasticity for the combination of pulse width and wavelength versus displacement into the surface of the laser processed edge. 3 is an SEM image showing the sidewall surface of a stent processed with various combinations of pulse width and wavelength. 3 is an SEM image showing the sidewall surface of a stent processed with various combinations of pulse width and wavelength. 3 is an SEM image showing the sidewall surface of a stent processed with various combinations of pulse width and wavelength. 3 is an SEM image showing the sidewall surface of a stent processed with various combinations of pulse width and wavelength.

  Embodiments of the invention relate to a method of laser processing a polymer substrate to produce a stent. More specifically, these embodiments provide a laser system and parameters that reduce or eliminate the adverse effects of the laser on the polymer and maximize the preservation of the functional properties of the material, such as the surface and bulk properties of the polymer material. And selecting and implementing.

  In general, a stent can have virtually any structural pattern that is compatible with the body lumen to be mounted. Typically, a stent is comprised of a circumferential ring and a pattern or network of longitudinally extending strut or bar arm interconnecting components. In general, the struts are arranged in a pattern, and these patterns are designed to contact the lumen wall of the blood vessel and maintain vascular patency.

  An exemplary structure of a stent is shown in FIG. FIG. 1 shows a stent 10 composed of struts 12. Stent 10 has interconnected cylindrical rings 14 connected by linking struts or links 16. The embodiments disclosed herein are not limited to making a stent or the stent pattern shown in FIG. These embodiments are readily applicable to other stent patterns and other devices. Variations in the structure of the pattern are virtually unlimited. The outer diameter of the produced stent (before crimping and deployment) can be between 0.2 and 5.0 mm. For coronary arteries, the diameter of the stent produced is 2.5-5 mm. The length of the stent can be between about 6-12 mm or more depending on the application.

  While this embodiment is particularly concerned with laser processing of polymer substrates to form stents, these methods involve other materials such as metals and ceramics, and composite materials composed of combinations of polymers, metals, and ceramics. Applicable to.

  The polymer can be biostable, bioabsorbable, biodegradable, or bioerodable. Biostability refers to a polymer that is not biodegradable. The terms biodegradable, bioabsorbable, and bioerodible, and the terms degradation, erosion, and absorption are used interchangeably and can be completely eroded or absorbed when exposed to body fluids such as blood, It also refers to a polymer that can be gradually reabsorbed, absorbed, and / or eliminated by the body. Furthermore, a drug-loaded stent can be made by coating the surface of the stent with an active substance or drug, or a biodegradable polymer carrier containing the active substance or drug. The drug coating is typically applied to the stent body or scaffolding after being formed by laser processing. This coating is typically much thinner than scaffolding struts, for example, the coating can be 1-5 microns in thickness, while struts are typically thick Is thicker than 100 microns, for example 140 to 160 microns thick.

  A mountable medical device such as a stent can be made by laser processing a structure or substrate to form a device. Material is removed from selected areas of the structure, resulting in the structure of the device. Specifically, a stent can be produced by processing a thin-walled tubular member with a laser. Selected regions of the tube material can be removed by laser machining to obtain a stent having a desired pattern. Specifically, the laser beam can be scanned over the surface of the tube material, or the tube material can be translated and rotated under the beam, removing trenches or kerfs that extend across the entire wall of the tube material. Is done. When the start point and the end point of the kerf coincide with each other, the region surrounded by the kerf falls or is removed by the assist gas. FIG. 2 illustrates one embodiment of a portion of a machine controlled system for laser machining a tube. In FIG. 2, the tube 200 is disposed within a rotatable collet fitting 204 of a machine-controlled device 208 for positioning the tube material 200 relative to the laser 212. In accordance with the machine encoding instructions, the tube 200 is rotated and moved axially relative to a laser 212 that is also mechanically controlled. The laser selectively removes material from the tube material and a pattern is cut into the tube. Thus, the tube is cut into a separate pattern of finished stents.

  FIG. 3 shows an enlarged view of the laser beam 408 interacting with the tube 414. Laser beam 408 is focused onto tube 414 by focusing lens 338. Tube 414 is supported at one end by a controlled rotating collet 337 and at the other end by an optional tube support pin 339. A coaxial gas jet assembly 340 directs a cold gas jet or stream 342 that escapes through a nozzle 344, which cools the processed surface as the beam cuts and cuts the substrate. The gas stream also helps remove debris from the kerf and cool the area near the beam. The gas inlet is indicated by arrow 354. A coaxial gas jet nozzle 344 is centered around the focused beam 352. In some embodiments, the supplied cooling gas pressure is between 30 psi and 150 psi. An exemplary flow rate of the cooling gas is between 2 and 100 scfh. Exemplary cooling or process gases include helium, argon, nitrogen, oxygen, or mixtures of these gases.

  Biodegradable polymers that can be suitable for stent scaffolding applications include semi-crystalline polymers. Specifically, these polymers include polymers having a glass transition temperature (Tg) above the human body temperature of about 37 ° C. Polymeric substrates include, but are not limited to, poly (L-lactide) (PLLA), polymandelide (PM), poly (DL-lactide) (PDLLA), polyglycolide (PGA), and poly (L-lactide-co- It can be made in whole or in part from a single biodegradable polymer or combination of biodegradable polymers, including glycolide) (PLGA). For PLGA, it includes copolymers containing various molar ratios of L-lactide to glycolide, such as 90:10, 75:25, 50:50, 25:75, and 10:90.

  Several properties of the stent are essential to performing its function, including radial strength and fracture resistance or elongation at break. For example, when a stent is crimped and deployed, the stent is subjected to significant stress / strain in the local area, so proper fracture resistance is required and is very important. While the stent pattern bend or inner or concave region 20 within the crown 18 shown in FIG. 1 is subject to high compressive stress and strain when the stent is crimped, the outer or convex region 22 of the crown 18 is When the stent is deployed it experiences high compressive stresses and strains. Therefore, the stent during crimping and deployment is very prone to cracking. Such cracks can lead to a loss of radial strength and potentially premature and / or serious failure of the stent. It is therefore essential that the mechanical properties of the polymer stent can be maintained throughout the laser processing process.

  In order to provide a stent with the desired properties, additional process steps can be introduced into the preform, i.e., the tube material, prior to its laser treatment to increase radial strength and puncture resistance. For example, fracture toughness can be greatly improved by controlling the size of the crystalline region and by optimizing the optimal amorphous / crystalline ratio for the semicrystalline polymer, and the radial strength can also be It can be seen that the polymer chain is preferentially arranged in the hoop direction. These desired microstructural properties can be tuned through a heated radial expansion step for the polymer that is higher than the Tg of the polymer. For example, for PLLA, the range of 65-120 ° C is preferred. Conversely, local heating of the substrate during laser cutting can cause the desired microstructural properties to be altered, or damage to local areas, thereby losing the benefits provided by tube processing. Or may decrease.

Laser beam processing is one of the most advanced non-contact processing techniques used in microfabrication and nanofabrication to meet the demands of advanced engineering materials, stringent requirements, complex shapes, and unique sizes. One. The present invention relates to a laser having a pulse width in the range of picoseconds (= 10 ⁻¹² ) (“picosecond” laser) and a laser having a pulse width in the range of femtoseconds (= 10 ⁻¹⁵ ). “Pulse width” refers to the duration of a light pulse with respect to time. The duration can be defined in several ways. Specifically, the pulse duration can be defined as the full width at half maximum (FWHM) of optical power with respect to time. Picosecond and femtosecond lasers offer unique advantages for removing accurate amounts of material with minimal thermal damage to surrounding materials. In general, picosecond lasers have a pulse width of less than about 10 ps, and femtosecond lasers have a pulse width between 10 fs and 800 fs.

  The two basic mechanisms used for laser ablation are thought to be the photothermal mechanism and the photochemical mechanism. In the photothermal mechanism, the material is ablated by melting and vaporization, while in the photochemical mechanism, the light energy of light is used to directly break the chemical bonds of the polymer. When the chemical bonds between the atoms and molecules of the substrate are broken, gas species are formed and these gas species are removed from the substrate.

  Laser ablation of material from the substrate can be performed by a thermal mechanism, a non-thermal mechanism, or a combination of both. For example, longer pulse lasers remove material from the surface primarily by thermal mechanisms. In the thermal mechanism, the absorbed laser energy raises the temperature at and near the absorption site and the material is removed by conventional melting or vaporization. The disadvantage of machining with this mechanism is that thermal damage to the uncut material occurs. Such damage includes melting at the work edge and thermal diffusion into a region or zone of material, which alters the properties of the substrate within that zone, causing cutting quality problems.

  Lasers with femtosecond pulse durations have recently been used to ablate material because the pulse duration is shorter than the typical thermalization characteristic time (ie, time to achieve thermal equilibrium) which is a few picoseconds. It is particularly important. Because the thermal diffusion depth is very small, it is considered completely or almost completely a non-thermal mechanism. Picosecond lasers remove material by a non-thermal mechanism for the most part, but also use some thermal mechanism for some materials, which is sufficient to cause some thermal damage to the substrate. .

  More specifically, the non-thermal mechanism involves optical breakdown in the target material that causes material removal. During photo breakdown of materials, a very high free electron density, ie plasma, is generated by mechanisms such as multiphoton absorption and avalanche ionization. When optical breakdown occurs, the target material is converted directly from its original solid state to a fully ionized plasma on a time scale that is too short to establish thermal equilibrium in the target material lattice. Therefore, there is little heat conduction beyond the area to be removed. As a result, there is little thermal stress or thermal shock to the material beyond about 1 micron from the laser processed surface.

  However, in the prior art, it is believed that it is not known whether non-thermal or photochemical mechanisms can cause damage to the uncut substrate that adversely affects stent performance, particularly fracture resistance. Specifically, it is not clear that optical breakdown propagates or penetrates beyond the laser processed edge surface and causes damage within the substrate. It is also not known whether such damage can affect properties that affect stent performance. In addition, the existing relationships of such laser parameters, such as laser beam wavelength, pulse energy, and pulse width when ablating polymer material, to such potential damage are unknown.

  The inventors have investigated that laser ablation of PLLA material by a non-thermal mechanism by using a femtosecond laser causes little thermal damage to the cutting surface, but the photochemical action is adjacent to the laser processed edge surface. It was found that damage (voids and cracks) was caused under the surface area. These damages reliably reduce stent performance. This is therefore important in selecting a laser system and its processing parameters to produce a polymer stent with minimal damage to stent properties such as fracture resistance.

  Embodiments of the present invention include defining process parameters for a picosecond pulsed laser to fabricate a polymer stent while preserving mechanical properties such as radial strength, elongation at break, or fracture resistance. Controlled laser parameters include the pulse width and wavelength of the laser energy. The use of such parameters in laser processing can minimize damage resulting from thermal and non-thermal mechanisms. In addition, embodiments include implementing a laser system and its parameters for laser processing used to fabricate PLLA and PLGA based substrates with minimal damage.

  We understand that the magnitude and depth of damage caused by photochemical mechanisms can be controlled by laser parameters such as pulse width and wavelength. In general, we understand that there is a tradeoff between damage caused by thermal and non-thermal mechanisms when laser processing polymers. As discussed below, both thermal and non-thermal mechanisms (ie, photochemistry) cause damage to the substrate. Although the characteristics of the damage are different, both can adversely affect stent performance. Laser parameters (eg, pulse width and wavelength) can be adjusted to reduce photochemistry, but somewhat increase the thermal mechanism.

  To illustrate the damage caused by shorter laser pulses, PLLA stent samples were fabricated by using femtosecond and picosecond range laser systems. Void formation can be seen in the bulk of the stent strut (below the surface). FIG. 4B shows that the voids are distributed over the surface area of the cutting edge.

  It is believed that high laser energy at the cutting edge converts a portion of the internal solid mass into a volume of gas that causes the formation of voids or bubbles. Voids are dispersed within a region extending from the cutting edge to a given depth, and dissipate beyond this depth. Such voids can act as stress concentrations that facilitate fracture formation when the stent struts are stressed, thus reducing fracture resistance and reducing fracture elongation. In addition to voids, cracks in the same area are also commonly observed in these samples.

  FIG. 4A shows a portion of a strut or structural element 500 formed by laser machining a substrate. The strut 500 has an external or external surface (eg, before processing the external surface of the tube) and a laser processed edge surface 508. The depth into the substrate perpendicular to the edge surface 508 is indicated by arrow 510. FIG. 4B shows a cross-section 512 of a portion of the strut 500 perpendicular to the processed edge surface 500 along AA. As shown by this cross section, the substrate has a void or bubble 516 having a diameter Dv and extending to a depth Dp.

  In addition, the modulus of elasticity for the stent surface material was measured (as a function of the distance from the processed edge surface for the stent). The change in modulus as a function of distance from the processed edge surface was observed. As shown in FIG. 9, the variation decreases as the distance increases into the surface and tends to converge to the elastic modulus of the virgin polymer unaffected by the laser energy. Additional tests were conducted on the mechanical properties of the altered region, including critical strength, elongation at break, modulus of elasticity, and maximum load, and these mechanical outputs were adversely affected by void formation.

  The size of the void and the depth at which the void exists depends on the laser parameters such as wavelength and pulse width. The void size can be less than 1 micron, 1-2 microns, 2-5 microns, or greater than 5 microns. Cracks or voids may be present at 2 microns or less, 5, 10, 15, 20, or 30 microns. In general, the mechanical properties of polymer stents are affected by the formation of these voids and cracks, which can be mitigated by proper selection of laser parameters such as wavelength and pulse width.

  Furthermore, thermal damage and non-thermal damage to the substrate material depends on the laser energy for a given pulse width for at least two reasons. First, the optical penetration of laser energy varies with wavelength. Second, the absorption coefficient, more generally the degree of absorption of laser energy by the laser treated polymer, varies with wavelength. In general, the lower the absorption coefficient in a given range of wavelengths, the greater the thermal effect on the substrate. For example, the absorption coefficient of PLLA increases from a negligible value at 800 nm and reaches a maximum at about 300-320 nm. Thus, as the wavelength decreases from about 800 nm to about 300-320 nm, photochemical removal increases and thermal removal decreases. Thus, the relative amount of non-thermal or thermal ablation also depends on the wavelength of the laser. This is because the polymer absorption coefficient depends on the wavelength of the laser.

  Thus, the relative amount of non-thermal / thermal ablation depends on both pulse width and wavelength. Furthermore, the degree of photochemical damage and thermal damage are both dependent on the pulse width.

  As indicated above, the present invention provides parameters such as pulse width and wavelength to preserve or maintain the mechanical properties of the stent and to reduce or minimize damage to the uncut portions of the substrate. Including adjusting. This tuning can include selecting one or more wavelengths between ultraviolet (10-400 nm) and infrared (greater than 700 nm) and one or more pulse widths. In some embodiments, the wavelength is selected such that the polymer has an absorption coefficient that is less than the maximum absorbance of the polymer at that wavelength, for example, the absorbance is 5-10%, 10-20%, 20-40% of the maximum absorbance. 40-60%.

  In some embodiments, the pulse width is adjusted to avoid excessive melting for one or more wavelengths, even if the level of cooling is appropriate during processing. Excessive melting may correspond to a molten material thickness greater than 0.25 microns, 0.5 microns, or greater than 1 micron.

For a given pulse width and wavelength, for example, the average laser to provide a sufficiently high flux (energy per pulse / beam spot size) so that the beam cuts the substrate across the entire wall of the polymer tube material. Power or power (energy per pulse x repetition rate) and repetition rate are selected. The beam spot size is typically 10-20 microns, but can be less than 10 microns or greater than 20 microns, depending on the application. Pulse energy and flux (based on 10 micron spot size) for laser cutting the polymer can be 4 to 200 μJ and 0.5 to 200 J / cm ² , respectively. The average power per pulse of the beam can be 0.5-4W or greater than 4W. In a narrower range, the power is 0.5 to 1 W, 1 to 1.5 W, 1.5 to 1.8 W, 1.8 to 2 W, 2 to 2.2 W, 2.2 to 2.5 W, 2. It can be set to 5-2.8W, 2.8-3W, 3-3.2W, 3.2-3.5W, 3.5-3.8W, 3.8-4W. For a 10 ps pulse width laser, the repetition rate can be 25-100 kHz, 25-50 kHz, 50-60 kHz, 60-80 kHz, or 80-100 kHz.

  Further, to reduce or minimize thermal effects (eg, melting at the surface of the cut) and to maximize cutting speed, the repetition rate and cooling gas flow rate (eg, unit SCFH He) is Adjusted or selected. In general, the repetition rate and the cutting speed are directly proportional. That is, the faster the repetition rate, the faster the cutting speed and the shorter the process time per stent. However, as the repetition rate increases, the thermal effect tends to increase. Increasing the cooling gas flow rate can mitigate the thermal effects from the increased repetition rate, allowing for higher repetition rates and hence cutting speeds. Thus, the repetition rate and cooling gas flow are selected to obtain the fastest process time with acceptable thermal effects.

  Laser processing in the picosecond range, for example 1-12 ps or higher, causes thermal effects, but these thermal effects are more effectively controlled than photochemical action, so it may be advantageous to work within this range. As indicated above, the thermal effect can be mitigated by cooling the substrate with a cooling gas as the substrate is processed.

  In addition, or alternatively, minimizing damage may correspond to minimizing the thickness of the altered region adjacent to the laser processed edge, including voids. For example, the thickness of this region can be less than 2 microns, less than 5 microns, less than 20 microns, or less than 30 microns. The void area can be 1-2 microns, 2-5 microns, 2-10 microns, 2-20 microns, or 5-10 microns. In addition, or in the alternative, minimizing damage can help minimize the variation in the modulus of the cut stent with distance within the damaged area. These parameters can be adjusted to achieve fast elastic modulus convergence towards the undamaged polymer elastic modulus. The modulus may converge at less than 4 microns, less than 8 microns, or less than 20 microns on the processed edge surface. The elastic modulus may converge between 1 to 4 microns, 4 to 8 microns, or 8 to 20 microns on the processed edge surface. The elastic modulus may converge from the processed edge surface below 4 microns, below 8 microns, below 15 microns, or below 20 microns.

  In addition, or alternatively, to minimize damage to the most desirable mechanical properties of the damaged area, such as critical strength, elongation at break, modulus of elasticity, or maximum load, by adjusting the wavelength and pulse width Can do. Most desirable is to have the cut polymer substrate maintain the same properties prior to its laser processing.

Further, embodiments include implementing laser systems and parameters when laser processing PLLA substrates to produce stents. In such an embodiment, the laser wavelength may be in the visible light spectrum of 390-800 nm. In some embodiments, the laser wavelength is in the green spectrum or 496-570 nm, or in a narrower range is 532 nm or 515 nm. Depending on the embodiment, the pulse width can be 0.8 ps or less, 0.8-1 ps, 1-5 ps, 5-10 ps, 10-12 ps, 12-15 ps, or greater than 15 ps. As disclosed in the Examples, the inventors have used a 10 ps laser and ablation at a wavelength of 532 nm and a repetition rate of 80 kHz, for PLLA stents compared to other combinations of lower pulse width and wavelength. It has been demonstrated that damage is minimized. Within the green wavelength range, or in particular at 532 nm, the repetition rate of the laser beam can be 25-100 kHz, or in the narrower range 25-40 kHz, 40-80 kHz, or 80-100 kHz. Table 1 provides the laser parameter ranges used in this study.

  As used herein, the following definitions apply:

  Unless otherwise specified, all ranges include endpoints and any value within an endpoint.

  The “radial strength” of a stent is defined as the pressure at which the stent undergoes irreversible deformation.

  “Stress” refers to the force per unit area, as in the case of a force acting through a small area in a plane. The stress can be divided into components called normal stress and shear stress, which are perpendicular and parallel to the plane, respectively. True stress refers to stress where force and area are measured simultaneously. Conventional stress applied in tensile and compression tests is the force divided by the original gauge length.

  “Maximum load” or ultimate load is the absolute maximum load (force) that a structure can withstand without failure.

  “Strength” refers to the maximum stress along an axis that a material will withstand before failure. The limit strength is calculated from the maximum load applied during the test divided by the original cross-sectional area.

  “Elastic modulus” is defined as the ratio of components of stress, or the force per unit area applied to a material divided by the strain along the axis of the applied force due to the applied force. be able to. The elastic modulus is the initial slope of the “stress-strain curve” and is therefore determined by the straight hook region of the curve. For example, some materials have both tensile and compressive moduli.

  “Strain” refers to the amount of elongation or compression that occurs in a material at a given stress or load.

  “Elongation” can be defined as the increase in length of a material that occurs when stressed. Generally expressed as a percentage of the original length.

  “Elongation at break” is the strain applied to a sample when it breaks. Usually expressed as a percentage.

  “Glass transition temperature” Tg is the temperature at which the amorphous region of the polymer changes from a brittle glass state to a solid deformable or ductile state at atmospheric pressure. In other words, Tg corresponds to the temperature at which segment motion in the polymer chain begins. The Tg of a given polymer is dependent on the heating rate and can be affected by the thermal history of the polymer. Furthermore, the chemical structure of the polymer greatly affects the glass transition by affecting mobility.

  The following examples and experimental data are for illustrative purposes only and are in no way intended to limit the invention. The following examples are provided to aid in understanding the present invention, but it should be understood that the present invention is not limited to the specific materials or procedures of the examples.

  The following set of examples describes the results of laser processing a stent from PLLA tube material for seven different parameter combinations of pulse width and wavelength. The PLLA tube material was formed from 100% PLLA resin by extrusion. The dimensions of the tubing, i.e. the extrusion dimensions, are outer diameter (OD) = 0.0066 inch and inner diameter (ID) = 0.0025 inch. Extruded PLLA tubes were conventionally radially expanded according to the method described in, for example, US Patent Application Publication No. 12 / 554,589, which is incorporated herein by reference. The target percent radial extension (% RE) was 400%. Here,% RE is transformed as 100% × (expanded tube inner diameter / tube original inner diameter-1).

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^５カリフォルニア州ペタルマのＲａｙｄｉａｎｃｅ Seven different parameter combinations of pulse width and wavelength are listed in Table 2.

¹ of Fremont, California Trumpf Laser Technology
² Germany, Lumera Laser GmbH of Kaiserslautern
³ Coherent, Inc. Santa Clara, California
⁴ IMRA in Fremont, California
⁵ Raydience in Petaluma, California

Additional laser parameters for these tests are given in Table 3. Velocity is the translational speed of the beam across the surface of the substrate. The pass is the number of times the cutting pattern is repeated. The processing time is the time required to cut the entire stent. Coherent's Libra (set number 5) corresponds to 800 nm in Table 3. The pico (532 nm) laser corresponds to set number 2. The laser in Table 3 is a fixed wavelength laser.

The properties of the surface area at the processed edge were evaluated using several test techniques. The techniques and properties obtained from these tests are summarized in Table 4. The test was performed on the samples using a cryoultramicrotome, nanoindentation, and tensile testing.

  The cryo-ultramicrotome refers to a technique of cutting ultrathin sections for microscopy. Thin sections of the processed stent perpendicular to the processed edges were cut and examined with a scanning electron microscope (SEM). The depth and size of voids and cracks were compared for different sets. Nanoindentation was used to measure the elastic modulus of the substrate as a function of distance from the processed edge. A tensile test was used to measure the mechanical properties of the surface area from the “dogbone” structure cut along the processed edge.

Cryogenic Ultramicrotome Results A summary of the cryoultramicrotome technique results for the four combinations of pulse width and wavelength is given in Table 5, and FIGS. 5-8 are adjacent to the processed surface for each combination of Table 5. An SEM image of the surface area is shown. For set number 5, significant photochemical action is indicated by bubbles extending to a depth of 20 μm. In set number 4, the bubble depth is significantly smaller than in set number 5, but the bubble size in set number 4 is larger, so it appears that a lower wavelength compared to set number 5 alleviated the photochemical action. Contrary to what was expected, the depth is greater for set number 4 because the penetration depth and wavelength are inversely proportional.

  For set number 1, cracks exist up to a depth of 30 μm. Set number 2 has the least damage among the four setups. The depth and size of bubble formation is the lowest from all test conditions.

A summary of nanoindentation results using five combinations of nanoindentation pulse width and wavelength is given in Table 6 and FIGS. The figure shows the elastic modulus versus displacement into the surface of the processed edge for 16 samples for each combination. Of the three combinations using a pulse width of 100 fs, the 800 nm sample (set number 5) shows the fastest convergence. Of the two combinations at 10 ps, the elastic modulus converges significantly faster at the wavelength of 532 nm. From a comparison of the 100 fs / 800 nm and 10 ps / 532 nm results, it can be seen that the fastest convergence of the overall modulus is provided by the 10 ps / wavelength 532 nm combination.

Tensile Test Tensile testing of the surface area of the processed edge was performed on samples processed with pulse width / wavelength combinations of 100 ps / 800 nm and 10 ps / 532 nm. The test was performed on a “dog bone” shaped sample made from the surface area of a laser processed substrate using a conventional tensile tester. Details for the 100 fs / 800 nm and 10 ps / 532 nm samples are given in Table 7 and Table 8, respectively.

Table 9 summarizes the results of the tensile test. The most notable difference in properties is the elongation at break. The 10 ps / 532 nm sample had an elongation over 2.5 times or about 60% greater than the 100 fs / 800 nm sample. Therefore, the former sample has significantly higher fracture resistance.

Summary of Material Results A summary of material results is provided in Table 10 for a set of samples. 10 ps / 532 sample has the best properties.

Processed Edge Quality Processed edge quality was investigated through observation of sidewall SEM images. One source of roughness is redeposition of vaporized substrate material. FIGS. 14-17 show SEM images showing stent sidewalls processed at 100 fs / 800 nm, 100 fs / 400 nm, 10 ps / 355 nm, and 10 ps / 532 nm, respectively. Sidewalls processed at the 10 ps / 532 nm combination have the smoothest processed edges.

  The above example shows that the wavelength of 532 nm provides the best balance of cutting speed with thermal effects and photochemical effects. This method can be applied to poly (L-lactide-co-glycolide) having any L-lactide to glycolide molar ratio between 0 and 1. The laser system and its settings can also be applied to cut metal substrates such as Ta—Nb—W and Co—Cr alloys.

While particular embodiments of the present invention have been shown and described, it will be apparent to those skilled in the art that changes and modifications can be made in its broader aspects without departing from the invention. Accordingly, the appended claims are to encompass within their scope all such changes and modifications as fall within the true spirit and scope of this invention.

Claims (12)

A method of laser processing a substrate to form a stent, comprising:
Providing a thin-walled polymer substrate;
The cutting through the thin wall, a laser beam having a predetermined pulse width and wavelength, the polymeric substrate of the thin wall and the laser processing, a step of forming a structural element having a machined edge surface, wherein A laser beam modifies the substrate in a surface region adjacent to the processed edge surface, and the modification is formed by converting a solid mass in the surface region into a gas by the laser energy of the laser beam. Including a void, a step,
Selecting the pulse width and wavelength so that the void exists at a depth of 2 μm or less;
Including methods. The method of claim 1, wherein the void has a diameter of less than 1 μm .   The method of claim 1, wherein the polymer is PLLA or PLGA.   The method of claim 1, wherein the polymer is PLLA, the wavelength is 532 nm, and the pulse width is 10 ps or less, 1 ps or more.   The method according to claim 1, wherein the repetition rate is 80 to 100 kHz. The pulse width and wavelength of the selected laser beam minimize damage to the surface region adjacent to the processed edge surface resulting from thermal and non-thermal ablation;
The method of claim 1, wherein the minimal damage is comprised of the void resulting from non-thermal ablation.   The method of claim 1, wherein the polymer is PLLA and the wavelength and the pulse width are adjusted in the green range and to 1-10 ps.   The method of claim 1, further directing a cooling gas in an area of the substrate that is processed by the laser beam. The method according to claim 1, wherein the surface region including the void has a thickness of 1 to 2 μm .   The method of claim 1, wherein the average power per pulse of the laser beam is 0.5-4 W. A plurality of interconnected structural elements formed by laser machining a thin-walled PLLA polymer tube with a laser beam directed to the outer surface of the tube, wherein the laser beam is formed on the wall of the tube A polymer stent body comprising a structural element that is cut through the wall of the polymer tube to remove selected portions to form the structural element having a side wall corresponding to a laser processed edge surface. And
The structural element comprises a plurality of bends or crowns that are subjected to stress and strain when the stent is crimped and deployed; the structural element connects the plurality of bends or crowns;
The surface region adjacent to the side walls of the plurality of bends or crowns and the structural element connecting the plurality of bends or multiple crowns may cause damage caused by the interaction of the laser beam with the tube. Have
The damage includes voids or bubbles dispersed within the surface region;
The polymer stent body, wherein the voids or bubbles extend from the laser processed edge surface to a depth of 1-2 μm , and the voids or bubbles have a diameter of less than 1 μm . Converging the injury, further comprising a variation of the elastic modulus of the polymer with distance from the laser machined edge surface, the variation of the elastic modulus, at a distance of 5 [mu] m or less from the laser machined edge surface The polymer stent body according to claim 11.

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